1. Field of the Invention
The present invention relates to a nuclear magnetic resonance imaging, and more particularly, to a nuclear magnetic resonance imaging suitable for imaging three dimensional physiological function information of an interior of a body to be examined at high speed.
2. Description of the Background Art
In recent years, among various types of medical diagnostic apparatuses developed, the nuclear magnetic resonance imaging (MRI) apparatus has been studied and developed very actively.
As well known, the nuclear magnetic resonance imaging is a method for imaging microscopic chemical and physical information of matters by utilizing the nuclear magnetic resonance phenomenon in which the energy of a radio frequency magnetic field rotating at a specific frequency can be resonantly absorbed by a group of nuclear spins having unique magnetic moments which are placed in a homogeneous static magnetic field.
This nuclear magnetic resonance imaging has attracted much attention because of its capability for imaging not just a physical shape information on a living body at high contrast, but also various other types of functional information such as a blood flow information, a microscopic magnetic field inhomogeneity information, a diffusion information, and a chemical shift information.
In particular, the study of the brain function imaging for imaging a change of the magnetic susceptibility of the hemoglobin in blood is very active recently.
Namely, as described in S. Ogawa et al.: "Oxygenation-Sensitive Contrast in Magnetic Resonance image of Rodent Brain at High Magnetic Fields", Magnetic Resonance in Medicine 14, pp. 68-78, 1990, it is known that, among the hemoglobin contained in blood of a living body, the oxyhemoglobin contained in abundance in the arterial blood is diamagnetic, while the deoxyhemoglobin mainly contained in the venous blood is paramagnetic.
Then, as described in R. M. Weisskoff et al.: "MRI Susceptometry: Image-Based Measurement of Absolute Susceptibility of MR contrast Agents and Human Blood", Magnetic Resonance in Medicine 24, pp. 375-383, 1992, it is also known that the diamagnetic oxyhemoglobin does not disturb a local magnetic field very much (magnetic susceptibility difference of 0.02 ppm with respect to living body tissues), but the paramagnetic deoxyhemoglobin has sufficiently large magnetic susceptibility difference with respect to surrounding tissues (magnetic susceptibility difference of 0.15 ppm with respect to living body tissues) to disturb the magnetic field so that the parameter T.sub.2. (which is a time constant parameter reflecting a lowering rate of the voxel signals based on an abrupt phase change of the nuclear spins caused by a microscopic magnetic field inhomogeneity change within a voxel) is going to be shortened.
Consequently, in peripheral veins within a region at which a brain function is activated where an amount of oxyhemoglobin is excessively increased, a local magnetic field inhomogeneity will be reduced compared with an inactive case and therefore voxel signals are going to be stronger. Hereafter, this effect will be referred as the BOLD (Blood Oxygen Level Dependent) effect.
Conventionally, an imaging scheme used for obtaining the brain function imaging has been the field echo imaging scheme, or the two dimensional echo planar imaging (2D-EPI) scheme illustrated in FIGS. 1 and 2. Here, FIG. 1 shows a typical pulse sequence in the 2D-EPI scheme, for a ease of acquiring the FID (Free Induction Decay) signals in which the BOLD effect is expected to be large. In the following, this case of acquiring the FID signals will be considered mostly. As can be seen from FIG. 1, this pulse sequence is characterized by a repeated inversion of the powerful reading gradient field pulse G.sub.r and a continuous application of the encoding gradient field short pulses G.sub.e. On the other hand, FIG. 2 shows a manner of data sampling (referred hereafter as k-trajectory) in the spatial frequency region (referred hereafter as k-space) of this scheme, in which the desired k-space data (in reading and encoding directions) are entirely acquired by a single excitation. In FIG. 2, an arrow indicates a direction of data sampling. In this case, the echo time will be optimally set to maximize the BOLD effect.
Now, in a practical realization of the brain function imaging, what is considered as one of the essential conditions is that it is "capable of comprehending functions of the entire brain three dimensionally, within a time (less than or equal to 1 second for example) in which an influence of a time change of a size or a position of the brain itself (in synchronization with the heart beat or the pulsation of cerebrospinal fluid) can be considerably reduced". In the following, whether this condition has been satisfied by some schemes proposed up until now will be considered in concrete details. Here, the basic conditions to be used in estimating the effect of each scheme will be set as follows: EQU X.sub.r =X.sub.e =192mm, X.sub.e,2 =96 mm, EQU .DELTA.X.sub.r =.DELTA.X.sub.e,2 =3mm, EQU N.sub.r =N.sub.e =64,N.sub.e,2 =32, EQU TE=50ms, EQU R.sub.DAT =1.0, and T1=787ms, T2=92ms,
where X.sub.r, X.sub.e, and X.sub.e,2 are imaging region (FOV: field of view) in reading, encoding, and second encoding (=slicing) directions, respectively, .DELTA.X.sub.r, .DELTA.X.sub.e, and .DELTA.X.sub.e,2 are spatial resolutions in reading, encoding, and second encoding (=slicing) directions, respectively, N.sub.r, N.sub.e, and N.sub.e,2 are numbers of image matrices in reading, encoding, and second encoding (=slicing) directions, respectively, TE is an echo time (with the BOLD effect taken into account). R.sub.DAT is an encoding direction data acquisition rate (which is equal to 1.0 in a full encoding method and 0.5 in a half encoding method). T1 is a longitudinal relaxation time, and T2 is a transverse relaxation time.
As a first conventional scheme, the 2D-EPI multi-slice scheme using a pulse sequence shown in FIG. 3 will be considered first. In this scheme, for three different strengths (5.8 mT/m, 13.5 mT/m, and 40 mT/m) of the reading gradient field G.sub.r (cases 1A, 1B, and 1C), an imaging time T.sub.scan, a relative image S/N, and levels of image degradation (image distortion, T2 image blurring and N/2 artifact) are as summarized in a table shown in FIG. 4. Here, the N/2 artifact is caused by the static magnetic field inhomogeneity, the system incompleteness, and the inversion of a sign of the reading magnetic field. Also, in FIG. 4, .DELTA.T.sub.sw,r is a read switching time (0.fwdarw. peak), M.sub.s is a number of multi-slices, DAT is a data acquisition time per one excitation, TR is an identical region excitation interval (repetition time), .alpha. is a flip angle, and TS is a sequence length per one slice.
As can be understood from FIG. 4, from case 1A to case 1C, the relative image S/N is lowered, while the various image degradation levels are improved and the imaging time is shortened, but the minimum value for the total imaging time to cover the entire FOV is limited by 2 sec., so that the above noted influence of motion cannot be ignored. Here, however, the imaging time per one slice TS is 62.5 ms to 100 ms. In order to achieve the total imaging time of less than or equal to 1 sec. for which the influence of motion can be reduced considerably, it is necessary to reduce a number of slices (i.e., the FOV) considerably. Moreover, the slice thickness of 3 mm raises a problem of a slice characteristic, and it becomes difficult to realize the isotropic resolution. Also, in case 1A, the relative image S/N is more than twice as much as that in case 1C, but the imaging time is even longer and the image degradation such as the N/2 artifact is even worse, so that this case 1A is practically intolerable.
Next, as a second conventional scheme, the 3D-EPI scheme with a pulse sequence shown in FIG. 5 will be considered. In this scheme, the the second encoding step is varied from one spin excitation to another by a constant value, and only two dimensional data in the reading and encoding directions are acquired by each excitation. This is a scheme which is equally popular as the first conventional scheme described above. In this scheme, for three different strengths (5.8 mT/m, 13.5 mT/m, and 40 mT/m) of the reading gradient field G.sub.r (cases 2A, 2B, and 2C), the imaging time T.sub.scan, the relative image S/N, and the levels of image degradation (image distortion, T2 image blurring, and N/2 artifact) are as summarized in a table shown in FIG. 6. Here, N.sub.e,2 indicates a number of sampling in the second encoding (3D) direction, and the flip angle .alpha. is approximated by the Ernst angle .alpha..sub.e in the FLASH scheme. Also, a reference value for the relative image S/N is set to be the value in case 1C of the first conventional scheme described above.
As can be understood from FIG. 6, from case 2A to case 2C, the relative image S/N is lowered, while the various image degradation levels are improved and the imaging time is shortened, but the minimum value for the total imaging time to cover the entire FOV is limited by 2 sec., so that the above noted influence of motion cannot be ignored. In order to achieve the total imaging time of less than or equal to 1 sec. for which the influence of motion can be reduced considerably, it is necessary to reduce the imaging region (FOV) considerably. These features are the same as in the first conventional scheme described above. What is different from the first conventional scheme is that it is impossible to obtain a partial image data unless the imaging is completed for the entire FOV, so that the influence of motion is more significant in this second conventional scheme. Here, however, the fact that the three dimensionally isotropic resolution can be obtained and the fact that the relative image S/N can be improved by the integrate effect are the merits over the first conventional scheme. Also, in case 2A, the relative image S/N is more than three times as much as that in case 2C due to the integrated effect in third dimensional direction, which is a further improvement over the first conventional scheme, but the imaging time is even longer and the image degradation such as the N/2 artifact is even worse, so that this case 2A is also practically intolerable just, like case 1A of the first conventional scheme.
Next, as a third conventional scheme, the scheme proposed by Mansfield, et al. (Japanese Patent Application Laid Open No. 2-131746 (1990)) with a pulse sequence shown in FIG. 7 and a k-trajectory shown in FIG. 8 will be considered. This scheme is basically a one shot 3D (volume) imaging=EVI, which is characterized by the fact that the second encoding gradient field pulse G.sub.e,2 is changed similarly as the first encoding gradient field G.sub.e during one excitation. Here, however, a number of echoes that can be acquired per one excitation is limited by the gradient field performance and the echo time, so that this scheme has an essential problem that if the same spatial resolution as in the other schemes is to be achieved, the FOV in the second encoding direction would be reduced considerably, whereas if the same FOV in the second encoding direction as in the other schemes is to be achieved, the spatial resolution would be degraded considerably.
In this scheme, for case 3C with 40 mT/m of the reading gradient field G.sub.r which is the only case having a sufficiently high gradient field performance for enabling the EVI, the imaging time T.sub.scan, the relative image S/N, and the levels of image degradation (image distortion, T2 image blurring, and N/2 artifact) are as summarized in a table shown in FIG. 9. Here, N.sub.echo is a total number of echoes acquired per one excitation. Also, the artifact due to T2 data discontinuity is an artifact caused by the step-wise change of the signal strength due to the T2 dissipation that occurs in each sampling in the second encoding direction, and this artifact becomes prominent when a number of encoding samplings is small compared with the data acquisition time per one excitation. Also, a reference value for the relative image S/N is set to be the value in case 1C of the first conventional scheme described above.
As can be understood from FIG. 9, the relative image S/X is twice as much as that in the first conventional scheme, and this is the integrated effect caused by the fact that the number of samplings in the second encoding direction is four times as much. The imaging time is as fast as 0.1 sec. because of the one shot imaging, but there is a fatal problem in that the FOV in the second encoding direction is considerably reduced to only 12 mm. In addition, there are also problems in that the artifact due to T2 data discontinuity becomes large, and that when the slice thickness is thin, it cannot take an advantage of the isotropy of the spatial resolution in the volume imaging because of the slice characteristic.
As discussed in concrete details above, it is shown that the first conventional scheme (2D-EPI multi-slice scheme) and the second conventional scheme (3D-EPI scheme) have an absolute limit in the shortening of the imaging time with respect to the desired resolution and the imaging region. Also, as for the third conventional scheme (EVI scheme), it is shown that the imaging time is sufficiently fast, but the imaging region size is considerably limited.
Consequently, up until now, there has been no proposition for a practical 3D brain function imaging scheme which satisfies the condition of being "capable of comprehending functions of the entire brain three dimensionally, within a time (less than or equal to 1 second for example) in which an influence of a time change of a size or a position of the brain itself can be considerably reduced". More generally, there has been no proposition for an ultra high speed 3D imaging scheme with sufficiently high image quality.